Virtual pet detector and quasi-pixelated readout scheme for pet

ABSTRACT

When designing detector arrays for diagnostic imaging devices, such as PET or SPECT devices, a virtual detector, or pixel, combines scintillator crystals with photodetectors in ratios that deviate from the conventional 1:1 ratio. For instance, multiple photodetectors can be glued to a single crystal to create a virtual pixel which can be software-based or hardware-based. Light energy and time stamp information for a gamma ray hit on the crystal can be calculated using a virtualizer processor or using a trigger line network and time-to-digital converter logic. Additionally or alternatively, multiple crystals can be associated with each of a plurality of photodetectors. A gamma ray hit on a specific crystal is then determined by a table lookup of adjacent photodetectors that register equal light intensities, and the crystal common to such photodetectors is identified as the location of the hit.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a divisional of U.S. patent application Ser. No.12/600,072 filed Nov. 13, 2009, which is the U.S. National Phaseapplication under 35 U.S.C. §371 of International Application No.PCT/IB2008/51663, filed Apr. 29, 2008, which claims the benefit of U.S.Provisional Application No. 60/938,282 filed May 16, 2007. Theseapplications are hereby incorporated by reference herein.

DESCRIPTION

The present application finds particular utility in positron emissiontomography (PET) and/or single photon emission computed tomography(SPECT) scanners in medical applications using pixels of different sizesor the like. However, it will be appreciated that the describedtechnique(s) may also find application in other types of scanningsystems and/or other medical applications.

Scintillator pixel size in PET is a primary factor determining thespatial resolution of the resulting image. Thus, depending on theapplication in mind, the scanner geometry and pixel size are optimizedto maximize the scanner performance and competitiveness. For instance,typical pixel size used in a full-body scanner is 4×4 mm², while in abrain or an animal scanner a pixel size of 2×2 mm² to 1×1 mm² may beused to increase the resolution. A 1:1 coupling between scintillatorpixels and photo-detectors is often used to optimize performance, andtranslates into a need for a custom-designed light detector for eachspecific application. This, however, implies significant changes to thelight detection chain as well. In conventional systems usingphotomultiplier tubes and Anger logic, this further means that the lightguide design and the size of the photomultiplier tubes must be adapted,thus leading to higher system development and manufacturing costs.

Having a larger number of detectors coupled with each scintillatorelement improves spatial resolution, and the plurality of detectors candetermine where in the scintillator the scintillation event occurs.However, timing resolution for determining time-of-flight measurementscan be compromised, since each of the multiple detectors only sees afraction of the light and may receive it with different delays. Bycontrast, coupling each scintillator with a corresponding detectoroptimizes the timing resolution, but reduces spatial resolution. Thesingle detector that receives all of the light from the crystal canprovide an accurate time stamp, but resolution is limited to the size ofthe scintillator.

Additionally, in conventional crystal-based PET systems, thescintillator crystals are coupled to the light detectors either on 1:1basis (pixelated readout) or through a light guide using Anger logic forcrystal identification. A drawback of pixelated readout is the vastnumber of channels and, as the crystal size gets smaller, the cost ofthe photodetectors. Anger logic-based systems, on the other hand, mapmany crystals onto few photodetectors, leading to smaller number ofchannels at the cost of increased pile-up and dead time and thus reducedcount rate capability.

Thus, there is an unmet need in the art for systems and methods thatfacilitate overcoming the deficiencies noted above.

In accordance with one aspect, a virtual pixel array for a diagnosticimaging system includes a virtual pixel comprising at least ascintillator crystal, a plurality of photodetectors optically coupled tothe scintillator crystal, which generate output signals in response toscintillations in the crystal, and a virtualizer that processes theoutput signals associated with a gamma ray hit on the scintillatorcrystal as detected by the plurality of photodetectors and calculates atime stamp for the gamma ray hit.

In accordance with another aspect, a method of method of calculating atime stamp for a virtual pixel includes receiving a gamma ray hit on ascintillator crystal of the virtual pixel, evaluating output signalsfrom each of a plurality of photodetectors optically coupled to thescintillator crystal to determine an energy and a photodetector timestamp for each photodetector associated with the gamma ray hit, andcalculating a total energy of the gamma ray hit by combining theenergies detected by the plurality of photodetectors associated with thegamma ray hit. The method further includes calculating a time stamp forthe gamma ray hit as a function of the photodetector time stampregistered by at least one photodetector in the plurality ofphotodetectors.

Yet another aspect relates to a detector array for a diagnostic imagingdevice, including a plurality of photodetectors arranged in an array,and a plurality of scintillator crystals arranged in an array andoptically coupled to the plurality of photodetectors, wherein thephotodetector array and the scintillator array being offset from eachother such that some of the scintillator crystals are coupled to adifferent number of photodetectors than other scintillator crystals. Thedetector array further includes a processor that identifies ascintillator crystal that has been hit by a gamma ray based on an outputsignal generated by one or more of the plurality of photodetectorsoptically coupled to the scintillator crystal hit by the gamma ray.

One advantage is that detector array design cost is reduced.

Another advantage resides in a universal detector array chip formultiple sizes of scanners.

Another advantage resides in improved compatibility between detectorarrays or scanners utilizing crystals of different sizes.

Still further advantages of the subject innovation will be appreciatedby those of ordinary skill in the art upon reading and understand thefollowing detailed description.

The innovation may take form in various components and arrangements ofcomponents, and in various steps and arrangements of steps. The drawingsare only for purposes of illustrating various aspects and are not to beconstrued as limiting the invention.

FIG. 1 illustrates a virtual pixel that is employed on a detector chipthat is configured to be optically coupled with any of a variety ofscintillator arrays.

FIG. 2 shows a hardware-based virtual pixel that is employed on auniversal chip in a manner similar to the software-based virtual pixeldescribed above.

FIG. 2A illustrates output signals from four photodetectors, responsiveto scintillation of a crystal with respective pulses A, B, C, and D.

FIG. 3 is an illustration of a hardware-based virtual pixel thatfacilitates using Anger logic to determine a time stamp for a gamma rayhit on the virtual pixel.

FIG. 4 illustrates an example of a quasi-pixelated readout scheme forPET imaging systems, maintaining the advantage of classical pixelatedreadout methods, including direct coupling of crystals to individualphotodetectors, while reducing the number of readout channels by afactor of four.

FIG. 5 illustrates a method of using a software-based virtual pixel tofacilitate combining smaller real pixels.

FIG. 6 illustrates a method of using a hardware-based virtual pixel tofacilitate combining smaller real pixels.

FIG. 7 illustrates a method of enabling a quasi-pixelated readout schemeand reducing readout channels for a detector comprising photodetectorsdirectly coupled to a number of scintillation crystals.

FIG. 8 illustrates a diagnostic imaging device such as may be employedin conjunction with one or more of the detectors, detector arrays, pixelarrays, virtual pixels, and/or methods described herein.

FIG. 1 illustrates a virtual pixel that is employed on a detector chipthat is configured to be optically coupled with any of a variety ofscintillator arrays (not shown). A “virtual” pixel can be a configurablecombination of photodetectors and scintillator crystals, designed to actor appear as a given standard type of detector, such as a 4×4 array of1×1 mm² photodetectors, or some other size array of a given size ofphotodetectors. The universal chip reduces manufacturing costs byenabling the same chip to be used on a variety of PET scanners. Althoughthe systems and methods described herein are presented largely withrespect to PET systems, it will be appreciated that such systems and/ormethods may be used in conjunction with SPECT systems, as well as otherradiation detection systems.

The redesign, manufacturing and testing of a new CMOS-basedphotodetector is a costly endeavor. To overcome this impediment, it isdesirable to design a digital silicon photomultiplier detector that issuited for all kinds of applications in PET and/or SPECT. This problemcan be surmounted by combining several small pixels into larger virtualpixels, which can be performed either in hardware, or in software. Forexample, four solid state detectors can be arranged in a square and canbe operated as four independent detectors, grouped together to operateas a single detector, etc. The virtual pixel facilitates introducingvirtualization to PET detectors based on a digital siliconphotomultiplier, which integrates the light sensing element and thereadout circuits onto the chip. Thus, the virtual pixel illustrated inFIG. 1 shows a software-based virtualization scheme, and subsequentfigures (showing hardware-based virtualization schemes) show additionalcircuits that facilitate connecting several smaller photodetectors 12into larger virtual pixels, which are matched to the application withoutsacrificing detector performance. As a result, a one-size-fits-alldetector is designed, thus saving costs on design and manufacturing andleading to improved readout performance compared to a monolithic pixelof the same size.

In one embodiment the virtual pixel comprises a scintillator crystal 10of a given size and multiple smaller photodetector pixels 12. Inaccordance with various aspects, an algorithm is employed by avirtualizer 14 (e.g. a processor) to combine the partial detector datainto a final “hit,” which describes a perceived gamma ray registrationevent at a portion of the virtual pixel crystal. Additionally, hardwarethat improves timing resolution and reduces data rate may be employed inconjunction with the virtual pixel.

The following example is presented to illustrate a scenario in which thevirtual pixels are useful. For instance, it may be desirable toimplement a photodetector with a smallest practical pixel size, such as1×1 mm² in the case of an animal scanner, and then glue larger crystalsto 2×2 or 4×4 photodetectors, if the detector is to be used in a brainor a human scanner. To further this example, either 4 or 16photodetector pixels can detect the photons emitted from the crystal andan external logic, implemented in a separate field programmable gatearray (FPGA) or the like, can combine the pixels to get the energy andthe time of the gamma hit. This, however, can lead to a reduced timingresolution, as the pixels see only ¼ or 1/16 of the light emitted by thecrystal, which can be particularly damaging as large pixels are used intime-of-flight PET scanners where the timing information is particularlyimportant. To build virtual pixels without compromising the timingresolution, the trigger lines of the photodetectors can be combinedtogether in a symmetrical way, as described below with regard to FIG. 2.

According to other aspects, software may be employed to permit a user toselect (e.g., via a pull-down menu or some other interface) virtualpixel size based on an application for which the virtual pixel is to beemployed in a detector array. Software may also be employed to selectactive trigger lines in a hard virtual pixel, which facilitatesmanipulating behavior of the virtual pixel.

FIG. 2 shows a hardware-based virtual pixel that is employed on auniversal chip in a manner similar to the software-based virtual pixeldescribed above. The virtual pixel includes trigger lines 22 thatconnect smaller photodetectors 12 comprised by the virtual pixel.Additionally, time-to-digital converters (TDC) 26 can be employed tomeasure the time of the first photon arriving at any of the combinedpixels. Moreover, a combination of the validation signal is employed tovalidate the resulting hit, as described in U.S. patent application Ser.No. 11/467,670, entitled “Digital Silicon Photomultiplier,” filed onAug. 28, 2006. The validation signal is used to determine if the eventis a real hit or if the acquisition has been started by a dark count. Itis used to discriminate real hits from noise, while the first triggerlevel that is used to stop the TDC is set just above the noise floor toget an optimal timing resolution. In the virtual pixel, a validationsignal is useful because the first cell that fires stops the TDC andstarts the acquisition. Therefore, in the case of a dark count, thevalidation prevents the pixel from executing a full acquisition and thendiscarding the data right afterwards. Doing so would increase the deadtime of the pixel, as any hit during that time would lead to anincorrect time stamp and also likely to an incorrect energy.

The validation works in the following manner: in a digital siliconphotomultiplier, a trigger network is hierarchically subdivided intovertical column lines that trigger a horizontal trigger line, which isthen connected to the TDC. The validation of a hit means that more thanone column trigger line indicates a hit for the hit to be validated. Forinstance, according to an example, if four columns exhibit activitywithin approximately 5 ns, then the event can be considered a real hit.Otherwise, the pixel can be quickly reset to get prepare for the nexthit. The probability of detecting 4 dark counts in a respective linewithin 5 ns is sufficiently low to make the foregoing validation schemehighly efficient. Thus, to validate a hit in a virtual pixel, at leastone of the photodetectors validates the hit. It will be appreciated thatother numbers of columns exhibiting activity (e.g., more than four, lessthan four, etc.), as well as other time periods (e.g., more than 5 ns,less than 5 ns, etc.) may be employed in conjunction with the variousvalidation schemes described herein, and that the foregoing example isillustrative in nature and not intended to limit the scope of thedescribed aspects.

One difference between “soft” and “hard” virtualization is that in thehard virtualization case, the photodetectors 12 included in the virtualpixel contribute directly to the time stamp of the hit, as their triggerlines are connected through a balanced network connected to anadditional TDC component 26. In some embodiments, the TDC componentincludes an accumulator as well, which sums partial photon counts fromthe respective photodetectors comprised by the virtual pixel. Thus, thephoton statistics remain substantially unchanged. Although a 2×2 mm²virtual pixel is shown in FIG. 2, four of the virtual pixels can be usedto implement a 4×4 mm² pixel, if desired, and so on. Aselector/multiplexer 28 selects which data to forward to one or moreoutput buffers (not shown). In one embodiment, software determines whichtrigger lines to activate in the hard virtual pixel in order tomanipulate pixel behavior.

In soft virtualization, the time stamp and energy of the final hit arecomputed by an algorithm, which can also be implemented in an externalFPGA for performance reasons, if desired. The algorithm uses the timestamps and energies from all pixels included in the virtual pixel, andcomputes the time stamp and the energy of the hit as if all pixels wereconnected together. While computing the total energy isstraight-forward, computing the time stamp can be more complicated.

Several ways of computing the time stamp of the combined hit arepossible. For instance, the time stamp can be computed in a way similarto that which is used in Anger-logic, as an energy-weighted sum of thepartial time-stamps. That is, as illustrated in FIG. 2A, each of thefour illustrated detectors responds to the scintillation with arespective pulse A, B, C, and D, whose beginning is indicative of timeand whose area is indicative of energy. However, this method involvesseveral multiply-add operations and a fixed-point division to beimplemented in hardware. Another method uses the time stamp of the hitwith the highest energy (e.g., C, in the example of FIG. 2A), assumingthis time stamp to be most accurate because of the photon statistics.Alternatively, the earliest time-stamp (e.g., A, in FIG. 2A) can be usedindependently of the energy of the partial hit—this method mimics thehard-wiring of the trigger lines of the virtual pixels. Although thereare some sources of error (e.g., quantum noise, TDC shifts, andnumerical errors) that may affect the time stamp more than in the hardvirtualization case, soft virtualization is relatively cheap and doesnot need any additional hardware, which can lead to a potentially highermanufacturing yield because of a lower number of gates.

FIG. 3 is an illustration of a hardware-based virtual pixel thatfacilitates using Anger logic to determine a time stamp for a gamma rayhit on the virtual pixel. For instance, a photodetector chip can beoptimized for a scanner with 2×2 mm² crystals, as shown in the figure.The same chip can be used in a full-body scanner, where 2×2 pixels arecombined to realize a 4×4 mm² virtual pixel with timing resolutionoptimized for time of flight (TOF). On the other hand, the same detectorcould also be connected to a light guide and 1×1 mm² crystals, forinstance in an animal scanner. Alternatively, a mapping scheme can alsobe used without a light guide. The circuitry implementing the Angerlogic can be placed on the chip itself, thus making the type of readouttransparent to the rest of the system. The orientation of thephotodetector pixels 12 comprised by the virtual pixel can be optimizedto minimize any “dead area” due to the electronics, as shown in FIG. 3.Additionally, space between pixels 12 can be utilized for processingcircuitry and the like.

To minimize the amount of circuitry used for the virtualization, someexisting components can be re-used (e.g., TDC, I/O buffers,accumulators, etc). The trigger network 22 forms a symmetrical balancedtree of buffers. Additionally, hit validation logic (not shown) isconnected in a similar way to facilitate separating triggers due to darkcounts from the real hits. At least one of the validation signals fromeach pixel can indicate a valid hit for the acquisition to continue. Thevalidation logic in the real pixels can be superseded by the validationlogic of the virtual pixel in order to accomplish this. Moreover, theacquisition sequence can be controlled by a virtual pixel state machine.

Additionally, the readout time of the virtual pixel is substantiallyidentical to the readout time of the small pixels making up the virtualpixel, compared to a monolithic large pixel where the readout timeincreases with pixel size. This is the result of the parallel readout ofthe smaller pixels, which can be done faster than in a large pixelhaving more lines to read.

It is to be understood that although various aspects described hereindistinguish between software-based virtual pixels and hardware-basedvirtual pixels, a combination of both hard and soft virtual pixels isintended to fall within the scope and spirit of this application. Forexample, a combination of separate hard and soft virtual pixels can beemployed in a single virtual pixel array. According to another example,a single virtual pixel can employ any or all of the features of a softvirtual pixel, such as described with regard to FIG. 1, as well as anyor all of the features of a hard virtual pixel, such as described withregard to FIGS. 2 and 3. Moreover, a virtual pixel array, whetheremploying soft or hard virtual pixels, can employ virtual pixels ofdifferent sizes, in order to localize a desired pixel size on a givenregion of interest or the like.

In another embodiment, a combination of virtual pixel sizes is used in asingle detector array. For instance, 1×1 virtual pixels can be employedabout a region of interest to maximize spatial resolution at thatlocation, while larger virtual pixels (e.g., 2×2, 4×4, etc.) may beemployed elsewhere. According to yet another embodiment, a 1×1 virtualpixel is employed for spatial resolution purposes, and a 2×2 or 4×4virtual pixel adjacent thereto is employed for timing resolutionpurposes. In this example, a pattern of alternating virtual pixel sizesis employed in the same detector array.

FIG. 4 illustrates an example of a quasi-pixelated readout scheme forPET or SPECT imaging systems, maintaining the advantage of classicalpixelated readout methods, including direct coupling of crystals toindividual photodetectors, while reducing the number of readout channelsby a factor of four. A detector array 50 is described, in which thenumber of photodetectors 52 and electronics channels used to achieve agiven level of sensitivity is reduced, thus allowing reduction in thecosts of the detector front-end. Additionally, direct coupling of thecrystals 54 to the photosensitive surface mitigates a need for a lightguide. Moreover, crystal pitch can be half the size of photodetectorpitch. As shown in FIG. 4, crystals 54 are connected directly to thesensitive surface of the photodetectors 52. The crystal size isapproximately ½ of the photodetector pitch, leading to a 4:1 mapping.The crystals are arranged such that one crystal delivers all light tothe center photodetector alone, 2 crystals share the light 1:1 with thedetectors adjacent to each side half, and 4 crystals share ¼ of thelight with the detectors adjacent at the corners. A simple lookuptable-based logic (e.g., discretized Anger logic or the like) can beused to identify the crystal by measuring the ratio of light sharedbetween the neighboring photodetectors.

The detector array 50 offers the advantage of allowing pixelated readoutwith only ¼ of the channels needed in conventional designs, thusreducing the costs of the electronics backend. Contrary to truepixelated readout, only 25% of the gamma ray hits lead to single-channeldead time. 50% of the hits lead to double-channel dead time, and in theremaining 25% of the hits, four channels are dead due to the lightsharing between neighboring photodetectors. The detector additionallyallows for the detection of Compton crosstalk when individualinteractions are separated by at least two crystals.

In order to combat X-ray fluorescence, which can lead to crystalmisidentification, the observation of a 3×3 photodetector field can beused as an input to a more refined discretized lookup table. In FIG. 4,for example, crystal 1 delivers 100% of its received light tophotodetector 1, crystals 2 and 3 each present 50% of their receivedlight with photodetector pairs 1 and 2, and 1 and 3, respectively.Crystal 4 distributes its received light equally to all fourphotodetectors, and so on.

According to another example, the mounting of the crystals 54 in astaggered, offset position relative to the photodetectors 52 improvesspatial resolution. For instance, if 100% of the light from a givencrystal is received by only one of the photodetectors, then a processor56 determines that the crystal is under the center of the photodetectorregistering the light. For example, the processor evaluates a lookuptable stored in memory 58 to identify the specific crystal transferringlight to the photodetector the position of which is known. In FIG. 4,when photodetector 1 registers a light transfer and no otherphotodetectors register a transference, the processor determines thatthe hit occurred on crystal 1, since crystal 1 is the only crystal thatcan transfer a full burst of light to photodetector 1.

If the amount of light received by two adjacent photodetectors is equal,then the processor similarly determines that the light is from thescintillation crystal that spans those two photodetectors. For instance,if photodetector 1 registers a transference of light, a lookup of thetable in memory 58, performed by the processor 56, will indicate thatthe hit was received at crystal 1, crystal 2, or crystal 3. Byevaluating whether other photodetectors registered an equal lighttransference, the processor can cross-index the photodetectors toisolate the precise crystal. In this example, if photodetector 2registers a light transference equal to that registered by photodetector1, then the processor determines that the hit occurred at crystal 2.Alternatively, if photodetector 3 registers a light burst equal to thatregistered by photodetector 1, then the processor determines thatcrystal 3 received the hit.

If the amount of the light received by four contiguous photodetectors isequal, then the scintillator crystal is determined to be positionedequally under the four detectors. In FIG. 4, if all four photodetectorsregister a substantially equal light transference, then the processordetermines that the hit was received at crystal 4. In this manner,relatively simple ratios can be used to resolve the spatial position ofthe scintillation with a resolution smaller than the size of thedetector.

It will be appreciated that good spatial resolution and good timingresolution are not mutually exclusive factors. For example, anembodiment can employ a pixel with one Geiger-mode avalanche photodiode(APD), which can facilitate achieving good timing resolution for anysize scintillator crystal of a given aspect ratio.

FIG. 5 illustrates a method 60 of using a software-based virtual pixelto facilitate combining smaller real pixels. At 62, a software-basedvirtual pixel is employed, such as the virtual pixel described withregard to FIG. 1. At 64, the time stamps and energies from all realpixels in the virtual pixel are evaluated. Total energy is computed at66, for instance by summing all energies registered on real pixels inthe virtual pixel. Time stamp for the hit registered on the virtualpixel is computed, at 68. Computing the time stamp is performedaccording to one or more different techniques.

For instance, according to one embodiment, the time stamp is computedusing an Anger logic technique, wherein the computed time stamp is theenergy-weighted sum of the partial time stamps. According to anotherembodiment, the time stamp computed at 68 is the time stamp of thehighest-energy hit, which is assumed to be the most accurate based onphoton statistics. According to yet another embodiment, the time stampcomputed at 68 is the earliest time stamp associated with the hit,independent of the energy of the partial hit. This embodiment mimics ahard-wired trigger line such as the trigger line described with regardto the hardware-based virtual pixel of FIGS. 2 and 3, as well as FIG. 6below. In this manner, virtual pixel cost is minimized because thesoftware-based virtual pixel does not need any additional hardware.

FIG. 6 illustrates a method 80 of using a hardware-based virtual pixelto facilitate combining smaller real pixels. At 82, a hardware-basedvirtual pixel is employed, which comprises a plurality of smaller realpixels, such as is described with regard to FIGS. 2 and 3, above. At 84,contributions of each real pixel to total energy and time stamp of thevirtual pixel are evaluated. At 86, the total energy registered at thevirtual pixels is calculated, such as by summing the energies of allreal pixels in the virtual pixel.

At 88, the time stamp for a gamma ray hit is determined using hardwiredtrigger lines and a TDC and component. For instance, since all pixelscomprised by the hard virtual pixel have trigger lines connected througha balanced network connected to a TDC, all pixels can contributedirectly to the time stamp of the hit.

According to another embodiment, Anger logic can be employed tofacilitate using a standard chip with a given crystal size for larger orsmaller scanning applications. For instance, a photodetector chip can beoptimized for a scanner with 2×2 mm² crystals, and same chip can be usedin a full-body scanner, if four 2×2 pixels are combined to realize a 4×4mm² virtual pixel with timing resolution optimized for TOF. Additionallyor alternatively, the same detector chip can be connected to a lightguide and 1×1 mm² crystals, for example in an animal scanner. Stillfurthermore, the 2×2 mm² chip can be coupled to 1×1 mm² crystals in amapping such as is described above with regard to FIG. 4. The circuitryimplementing the Anger logic can be placed on the same chip, thus makingthe type of readout transparent to the rest of the system. Theorientation of the photodetector pixels building a virtual pixel can beoptimized to minimize the dead area due to the electronics, as shown inFIG. 3, above.

FIG. 7 illustrates a method 100 of enabling a quasi-pixelated readoutscheme and reducing readout channels for a detector comprisingphotodetectors directly coupled to a number of scintillation crystals,such as detector 50 described above. In this example, the crystals havea pitch one-half the size of the pitch of the photodetectors to whichthey are coupled. According to one embodiment, At 102, a hit is detectedon a first photodetector. At 104, a determination is made regardingwhether a second photodetector, adjacent to the first photo detector,has registered a hit comprising a substantially equal amount of light.At 106, a table lookup is performed for all adjacent photodetectorsregistering substantially equal hits, in order to determine where thehit was received. At 108, a crystal that received the hit is identifiedbased on the photodetectors that registered the substantially equalhits.

According to an example, a first photodetector registers a hit at 102and no second photodetector registers a hit per the determination at104. In this scenario, a lookup of the table at 106 results in anidentification of a single crystal at 108, which is the crystal that iscoupled solely to the first photodetector and does not overlay any otherphotodetectors. According to another example, wherein two adjacentphotodetectors receive hits of substantially equal magnitude as detectedat 102 and 104, then the lookup at 106 will result in an identification,at 108, of the crystal that overlaps both of the photodetectors and noother photodetector. According to yet another example, fourphotodetectors register substantially equal hits as determined at 102and 104. In this case, the lookup at 106 will identify the crystal thatoverlaps a corner of all four photodetectors, at 108.

FIG. 8 illustrates a diagnostic imaging device 120 such as may beemployed in conjunction with one or more of the detectors, detectorarrays, pixel arrays, virtual pixels, and/or methods described herein.The diagnostic imaging device 120 includes a housing 122 and a subjectsupport 124. Enclosed within the housing 122 is a detector array 126.The detector array 126 includes a plurality of individual detectorelements 128. The array 126 is arranged so that detector elements 128are distributed evenly about an imaging region 130. The detector array126 can be a ring of detectors 128, multiple rings, or discrete flatpanels disposed opposing each other. Whatever the actual placement orarrangement of the detectors 128, it is preferable to arrange thedetectors such that each detector has a plurality of counterpartdetectors across the imaging region to facilitate coincidence detection.In positron emission tomography (PET), pairs of gamma rays are producedby a positron annihilation event in the imaging region and travel inopposite directions. These gamma rays are detected as pairs, with aslight delay (on the order of nanoseconds) between detections if onegamma ray travels farther to reach a detector than the other.

Before the PET scan commences, a subject is injected with aradiopharmaceutical. The radiopharmaceutical contains a radioactiveelement coupled to a tag molecule. The tag molecule is associated withthe region to be imaged, and tends to gather there through normal bodyprocesses. For example, rapidly multiplying cancer cells tend to expendabnormally high amounts of energy duplicating themselves. So, theradiopharmaceutical can be linked to a molecule, such as glucose that acell typically metabolizes to create energy, gather in such regions andappear as “hot spots” in the image. Other techniques monitor taggedmolecules flowing in the circulatory system.

For PET imaging the selected radioisotope emits positrons. The positroncan only move a very short distance (on the order of nanometers) beforeit is annihilated in an annihilation reaction that creates twooppositely directed gamma rays. The pair of gamma rays travel inopposite directions at the speed of light striking an opposing pair ofdetectors.

When a gamma ray strikes the detector array 126, a time signal isgenerated from a leading edge of the resultant electrical pulse. Atriggering processor 132 monitors each detector 128 for an energy spike,e.g., integrated area under the pulse, characteristic of the energy ofeach received gamma ray. The triggering processor 132 checks a clock 133and stamps each detected gamma ray with a time of leading edge receiptstamp. The time stamp is first used by an event verification processor134 to determine which gamma rays are a pair which defines a line ofresponse (LOR). Because gamma rays travel at the speed of light, ifdetected gamma rays arrive more than several nanoseconds apart, theyprobably were not generated by the same annihilation event and arediscarded. Timing is especially important in TOF-PET, as the minutedifference in substantially simultaneous events can be used to furtherlocalize the annihilation event along the LOR. As computer processorclock speeds become faster, the higher the accuracy with which an eventcan be localized along its LOR. In a SPECT camera, the LOR or trajectoryfor each detected gamma ray is determined by collimation.

LORs are stored in an event storage buffer 144, and a reconstructionprocessor 146 reconstructs the LORs into an image representation of thesubject using filtered backprojection or other appropriatereconstruction algorithm. The reconstruction can then be displayed for auser on a display device 148, printed, saved for later use, and thelike.

Having thus described the preferred embodiments, the invention is nowclaimed to be:
 1. A virtual pixel array for a diagnostic imaging system,including: a virtual pixel comprising at least one scintillator crystal;a plurality of photodetectors optically coupled to the at least onescintillator crystal, which generate output signals in response toscintillations in the crystal, wherein the plurality of photodetectorsincludes 4 photodetectors arranged in a 2×2 array; and a virtualizerthat processes the output signals associated with a gamma ray hit on thescintillator crystal as detected by the plurality of photodetectors andcalculates a time stamp for the gamma ray hit; a plurality ofscintillator crystals, including the at least one scintillator crystal,arranged in a rectangular grid; a plurality of the 2×2 arrays ofphotodetectors optically coupled to the scintillator crystals in anoffset relationship such that in each 2×2 array, one of thephotodetectors is optically coupled to only one scintillator crystal,one of the photodetectors is optically coupled to two of thescintillator crystals, and two of the photodetectors are opticallycoupled to four of the scintillator crystals.
 2. The virtual pixel arrayaccording to claim 1, wherein the length of each side of eachscintillator crystal is approximately ½ the length of each side of eachphotodetector, such that each photodetector is associated with ninescintillator crystals.
 3. A method of identifying a scintillator crystalin the virtual pixel array of claim 1, including: detecting lightregistration from a gamma ray at a first photodetector; determiningwhether at least a second photodetector, adjacent to the firstphotodetector, has registered an amount of light equal to the amount oflight registered at the first photodetector; performing a table lookupof photodetectors registering substantially equal amounts of light; andidentifying a scintillator crystal overlapping all photodetectorsregistering the substantially equal amount of light as the scintillatorcrystal that was hit by the gamma ray.
 4. The virtual pixel arrayaccording to claim 1, further including: one or more of: at least one4×4 array of photodetectors optically coupled to the scintillatorcrystals in an offset relationship; and at least one 1×1 array ofphotodetectors optically coupled to the scintillator crystals in anoffset relationship.
 5. A method of calculating a time stamp for avirtual pixel, including: arranging a plurality of scintillatorcrystals, including the at least one scintillator crystal, in arectangular grid; optically coupling a plurality of 2×2 arrays ofphotodetectors to the scintillator crystals in an offset relationshipsuch that in each 2×2 array, one of the photodetectors is opticallycoupled to only one scintillator crystal, one of the photodetectors isoptically coupled to two of the scintillator crystals, and two of thephotodetectors are optically coupled to four of the scintillatorcrystals receiving a gamma ray hit on at least one scintillator crystalof the virtual pixel; evaluating output signals from each of a pluralityof photodetectors optically coupled to the at least one scintillatorcrystal to determine an energy and a photodetector time stamp for eachphotodetector associated with the gamma ray hit; calculating a totalenergy of the gamma ray hit by combining the energies detected by theplurality of photodetectors associated with the gamma ray hit; andcalculating a time stamp for the gamma ray hit as a function of thephotodetector time stamp registered by at least one photodetector in theplurality of photodetectors.
 6. The method according to claim 5, whereinthe length of each side of each scintillator crystal is approximately ½the length of each side of each photodetector, such that eachphotodetector is associated with nine scintillator crystals.
 7. Themethod according to claim 5, further comprising: detecting lightregistration from a gamma ray at a first photodetector; determiningwhether at least a second photodetector, adjacent to the firstphotodetector, has registered an amount of light equal to the amount oflight registered at the first photodetector; performing a table lookupof photodetectors registering substantially equal amounts of light; andidentifying a scintillator crystal overlapping all photodetectorsregistering the substantially equal amount of light as the scintillatorcrystal that was hit by the gamma ray.
 8. The method according to claim5, further including: one or more of: optically coupling at least one4×4 array of photodetectors to the scintillator crystals in an offsetrelationship; and optically coupling at least one 1×1 array ofphotodetectors to the scintillator crystals in an offset relationship.9. A detector array for a diagnostic imaging device, including: aplurality of photodetectors arranged in an array; a plurality ofscintillator crystals arranged in an array and optically coupled to theplurality of photodetectors, the photodetector array and thescintillator array being offset from each other such that some of thescintillator crystals are coupled to a different number ofphotodetectors than other scintillator crystals; and a processor thatidentifies a scintillator crystal that has been hit by a gamma ray basedon an output signal generated by one or more of the plurality ofphotodetectors optically coupled to the scintillator crystal hit by thegamma ray.
 10. The detector array according to claim 9, wherein thelength of each side of each scintillator crystal is approximately ½ thelength of each side of each photodetector, such that each photodetectoris associated with nine scintillator crystals.
 11. The detector arrayaccording to claim 9, further including: a plurality of the scintillatorcrystals arranged in a rectangular grid; a plurality of 2×2 arrays ofphotodetectors optically coupled to the scintillator crystals in anoffset relationship such that in each 2×2 array, such that one of thephotodetectors is optically coupled to only one scintillator crystal,one of the photodetectors is optically coupled to two of thescintillator crystals, and two of the photodetectors are opticallycoupled to four of the scintillator crystals.
 12. A method ofidentifying a scintillator crystal in the detector array of claim 9,including: detecting light registration from a gamma ray at a firstphotodetector; determining whether at least a second photodetector,adjacent to the first photodetector, has registered an amount of lightequal to the amount of light registered at the first photodetector;performing a table lookup of photodetectors registering substantiallyequal amounts of light; and identifying a scintillator crystaloverlapping all photodetectors registering the substantially equalamount of light as the scintillator crystal that was hit by the gammaray.
 13. A method of identifying a gamma ray hit location on a detectorarray of claim 9, including: arranging a plurality of the scintillatorcrystals in a rectangular grid; optically coupling a plurality of the2×2 arrays of photodetectors to the scintillator crystals in an offsetrelationship such that in each 2×2 array, one of the photodetectors isoptically coupled to only one scintillator crystal, one of thephotodetectors is optically coupled to two of the scintillator crystals,and two of the photodetectors are optically coupled to four of thescintillator crystals.
 14. The detector array according to claim 9,further including: a plurality of the scintillator crystals arranged ina rectangular grid; a plurality of 4×4 arrays of photodetectorsoptically coupled to the scintillator crystals in an offsetrelationship.
 15. The detector array according to claim 14, furtherincluding: wherein the photodetectors are 1×1 mm² photodetectors. 16.The detector array according to claim 9, further including: a pluralityof the scintillator crystals arranged in a rectangular grid; and two ormore of: a 4×4 array of photodetectors optically coupled to thescintillator crystals in an offset relationship; a 2×2 array ofphotodetectors optically coupled to the scintillator crystals in anoffset relationship; and a 1×1 array of photodetectors optically coupledto the scintillator crystals in an offset relationship.
 17. The detectorarray according to claim 16, further including: wherein thephotodetectors are 1×1 mm² photodetectors.